The present invention relates generally to a magnetic resonance imaging (MRI) scanner and more particularly to a low acoustic noise MRI scanner.
MRI scanners, which are used in various fields such as medical diagnostics, typically create images based on the operation of a magnet, a gradient coil assembly, and a radiofrequency coil(s). The magnet creates a uniform main magnetic field that makes unpaired nuclear spins, such as hydrogen atomic nuclei, responsive to radiofrequency excitation via the process of nuclear magnetic resonance (NMR). The gradient coil assembly imposes a series of pulsed, spatial-gradient magnetic fields upon the main magnetic field to give each point in the imaging volume a spatial identity corresponding to its unique set of magnetic fields during an imaging pulse sequence. The radiofrequency coil applies an excitation rf (radiofrequency) pulse that temporarily creates an oscillating transverse nuclear magnetization in the sample. This sample magnetization is then detected by the excitation rf coil or other rf coils. The resulting electrical signals are used by the computer to create magnetic resonance images. Typically, there is a radiofrequency coil and a gradient coil assembly within the magnet.
Magnets for MRI scanners include superconductive-coil magnets, resistive-coil magnets, and permanent magnets. Known superconductive magnet designs include cylindrical magnets and open magnets. Cylindrical magnets typically have an axially-directed static magnetic field. In MRI systems based on cylindrical magnets, the radiofrequency coil, the gradient coil assembly and the magnet are generally annularly-cylindrically shaped and are generally coaxially aligned, wherein the gradient coil assembly circumferentially surrounds the radiofrequency coil and wherein the magnet circumferentially surrounds the gradient coil assembly. Open magnets typically employ two spaced-apart magnetic assemblies (magnet poles) with the imaging subject inserted into the space between the assemblies. This scanner geometry allows access by medical personnel for surgery or other medical procedures during MRI imaging. The open space also helps the patient overcome feelings of claustrophobia that may be experienced in a cylindrical magnet design.
A gradient coil assembly comprises a set of windings in a support structure that produce the desired gradient fields. Such an assembly for a human-size whole-body MRI scanner typically weighs about 1000 kg. The windings consist of wires or conductors formed by cutting or etching sheets of conducting material (e.g. copper) to form current paths to generate desired field patterns. The wires or conducting coils or plates are themselves typically held in place by fiberglass overwindings plus epoxy resin.
Generally, the various components of the MRI scanner represent sources and pathways of acoustic noise that can be objectionable to the patient being imaged and to the operator of the scanner. For example, the gradient coil assembly generates loud acoustic noises, which many medical patients find objectionable. The acoustic noises occur in the imaging region of the scanner as well as outside of the scanner. Known passive noise control techniques include locating the gradient coil assembly in a vacuum enclosure.
Large pulsed electrical currents, typically 200 A or more, with rise times and durations typically in the submillisecond to millisecond range, are applied to the windings. Because these windings are located in strong static magnetic fields (e.g., 1.5 T for a typical clinical imager to much higher values for research systems), the currents interact with the static field and strong Lorentz forces are exerted on different parts of the gradient coil assembly. These forces in turn compress, expand, bend or otherwise distort the gradient coil assembly. It will be readily understood by those skilled in the art that the frequencies of the acoustic noise so generated will be in the audio range. Typically there are strong components of noise from 50 Hz and below to several kHz at the upper end of the frequency range.
The magnet cryostat and other metallic parts of the magnet assembly are also sources of vibration and acoustic noise. For a cylindrical magnet, the actively shielded gradient winding generally comprises three layers, one for each Cartesian direction (x, y and z), each layer typically consisting of an inner, primary winding portion that creates a gradient field in the imaging region plus an outer, concentric shielding winding portion that substantially reduces the fields outside the gradient assembly. Some of the gradient-produced fields leak through or around the shielding. These leakage fields can induce eddy currents in the magnet metallic parts, for example, the cryostat inner bore. These eddy currents in turn produce Lorentz forces on the cryostat inner bore leading to mechanical motion of the cryostat inner bore and consequent acoustic noise according to WA Edelstein et al., Magnetic Resonance Imaging 20, 155-163, 2002.
The shielded gradients described above are constructed with inner gradient windings to produce gradient fields Gx, Gy and Gz of the main Bo field along axes x, y and z. They also have larger diameter shielding windings designed to substantially cancel external fields produce by the inner gradient windings so that the net fields outside the shielding windings are substantially zero. There are generally three separate inner windings, one for each of the directions x, y and z, and three separate outer, shielding windings, one for each of the directions x, y and z. The complete x-gradient winding consists of various parts, including inner and shielding windings, which are generally electrically connected to form a single electrical circuit. The configurations of y- and z-gradient windings are similar. Up to the present, the general practice has been to design the inner and shielding windings so that each axis has a single layer of inner windings and a single layer of shielding windings. Typically, the inner x-winding is electrically connected in series to the outer, shielding x-winding and the inner plus outer x-winding circuit is then powered by a single power supply. The same is true of the y- or z-winding.
FIG. 1A is a cross-sectional side view of a conventional actively shielded gradient. 304 is the metallic inner bore of the magnet cryostat. 20 is the inner gradient windings and 30 is the outer, shielding gradient windings. 20 and 30 are supported in 102, a solid, nonconducting annular support structure.
However, the active gradient shielding is not perfect, as some pulsed magnetic fields leak through and around the shielding windings, interact with the magnet or cryostat structure and cause eddy currents in those structures. As shown in Schenck et al., U.S. Pat. No. 4,617,516, 1986, the z-gradient windings are generally made using a wire wound in one direction from one end of the gradient to the center, and the wire reverses from the center outward. Substantial field can leak through the widely spaced turns in the center. For all windings, field can leak around the ends of the windings to interact with the cryostat or other metallic parts of the magnet structure. It has been shown that eddy currents in the metallic inner bore of the magnet cryostat produced Lorentz forces on the inner bore leading to vibrations of the inner bore and consequent acoustic noise (WA Edelstein et al., Magnetic Resonance Imaging 20, 155-163, 2002.)
Some recent designs to improve gradient active shielding produce increasingly complex gradient current patterns. The length of these windings must be limited to ensure efficient gradient operation. Compromises thereby incurred limit the effectiveness of the additional shielding achieved. The recent quest for shorter length and rapidly changing gradients has also tended to work against effective shielding.
The active gradient shields in cylindrical magnet MRI systems heretofore has been confined to a cylindrical surface disposed concentrically relative to the magnet. This arrangement is not the best configuration to produce the most effective shielding.
Passive shielding alone will not provide adequate shielding and eddy current control because of its finite time constants. Multiple layers of active gradient shielding will have limited effectiveness because active gradient shielding must have discrete current paths. We are proposing the use of passive gradient shielding in conjunction with active gradient shielding. One example of this approach is disclosed in Mulder et al., U.S. Pat. No. 6,326,788, 2001. FIG. 3 is a cross-sectional side view of their design. It is similar to the conventional actively shielded gradient shown in FIG. 1 with the addition of a passive shield 210 consisting of a conducting layer on the outer radius of the solid, nonconducting annular cylindrical support structure 102. The idea of passive shielding acting in combination with active shielding can be extended further with substantially improved efficacy.
What is needed is a method of drastically reducing gradient-induced eddy currents in the magnet cryostat and other magnet parts, and to mitigate consequent vibrations, in order to substantially alleviate one of the principal sources of acoustic noise in MRI scanners.